Interest in chemical and biochemical sensors has expanded rapidly in recent years. One of the more recent developments has been the attempt to develop immunosensors based on optical fibres, which are usually characterized by their high specificity and high sensitivity. These and other developments focused on an attempt to adapt the principles of solid-phase immunoassay to continuous monitoring of various substances e.g. drugs, hormones etc. in body fluids. Immunosensors developed for these purposes are usually based on the conventional competitive immunoassay reaction between an immobilized receptor present in a solid phase on the one hand and a mobile ligand (analyte) and a mobile labelled analogue of the latter present in a liquid phase on the other hand. Such a reaction may be described by the following equation: EQU Ab+Ag+Ag*.fwdarw.Ab-Ag+Ab-Ag*
In this equation Ab is the receptor and stands for an antibody and Ag and Ag* are the ligands and stand for an antigen analyte and a labelled derivative of the same antigen or of an analogue thereof, respectively.
According to the elementary immunoassay method for the determination of an analyte Ag in a test liquid, a sample of the tested liquid containing a known amount of labelled antigen Ag* and an unknown amount of the analyte Ag is introduced into a vessel holding a solid phase containing a given amount of an appropriate antibody Ab. Upon contact, the analyte antigen Ag present in the sample and the added labelled antigen Ag* compete for the limited amount of antibody Ab in the solid phase. After a certain incubation period the liquid and the solid phases are separated from each other, the solid phase is washed and the analyte concentration is determined by following the labelled component, either as Ag* in the liquid phase or as the conjugate Ab-Ag* in the solid phase.
In known solid-phase immunoassays only one component of the competitive recognition system (usually the antibody Ab) is immobilized on the sensing area, while the labelled antigen Ag* is introduced ad hoc into the liquid sample which is tested for the antigen Ag.
The above-described immunoassay method is static and unsuitable for the continuous determination of analyte in real or nearly real time. A further development in the field of immunoassay was achieved by the employment of probes which enable discrete measurements of analytes, although not in real time. In one version of these probes the sensing tip, with predetermined amounts of immobilized receptors, is dipped in a test liquid containing a known amount of labelled ligand and unknown amount of analyte ligand. From the relative amount of the labelled ligand which conjugated to the receptors, present on the sensing tip, the concentration of the analyte is determined. By a sequence of such discrete measurements the change in concentration of the analyte with time may be determined in close approximation.
In sensors in which a labelled antigen or a labelled receptor (one of them is an analogue to the analyte) is used, the concentration of the labelled species has to be kept constant in order to serve as a reliable reference for the continuously changing concentration of the analyte. Subject to this it is further required that the receptor-ligand affinity interaction be specific and reversible so that the system will responds continuously and reversably to changes of analyte concentration.
Various approaches have been proposed to solve the latter problem. Some of these proposals are based on a compromise in that a degree of reversibility is obtained by so selecting the receptor that the affinity between receptors and ligands is moderated. However, by its very nature such a compromise solution is not quite satisfactory as the sensing is less specific and it is liable to yield false results. Consequently, efforts are continuously made to find better solutions for the desired reversibility while keeping high specificity of the competitive recognition system.
Recently, some approaches, based mainly on encapsulated reagents, have been proposed (Refs. 1-5). Thus, for example, Mansouri and Schultz (Ref. 2) have developed an optical affinity glucose sensor, in which a labelled glucose analogue in the form of high molecular weight fluoresceindextran is entrapped within a dialysis fibre having an outer membrane permeable to glucose. A given amount of concanavalin-A is covalently attached to the inner lumen of the dialysis fibre, and the entrapped fluoresceinated dextran and permeating glucose compete for conjugation therewith. At equilibrium the level of free fluorescein in the hollow fibre lumen is measured via the optical fibre and is correlated to the concentration of glucose. The higher the glucose concentration, the higher the recorded fluorescence.
Similar biosensors have been proposed by Meadows and Schultz (Ref. 3) and by Anderson and Miller (Ref. 4). In these systems, which are based on fluorescence energy transfer, the specific receptor and the high molecular weight labelled ligand are enclosed within a dialysis membrane, permeable to the low molecular weight unlabelled ligand such as glucose or phenytoin.
A further system, based on potentiometric response, has been proposed by Bush and Rechnitz (Ref. 5). A monoclonal antibody Ab is trapped between two membranes while its antigen (DNP) is immobilized on one of the membranes.
In all the known biosensor systems according to Reference 1 to 5, one or more of the competitive reagents is enclosed within a membrane permeable to the low molecular weight analyte molecules, but not permeable to their labelled analogues. This is essential in order to meet the requirement for a constant level of labelled analyte analogues within the sensing zone. The result is that all these biosensing systems are suitable only for low molecular weight analytes.
A further disadvantage of the above prior art sensors is that since the analyte has to diffuse through the membrane, the response time is increased.
Recently, Davis and Leary (1989), Ref. 6, have developed a piezoelectric device for kinetic immunoassay which is based on the change of crystal frequency as a function of analyte concentration. Although the authors refer to this as a biosensor, within the definitions herein this piezoelectric device is a probe and not a sensor. Thus, although Davis's system does not require the use of a labelled antigen and consequently no semipermable membrane, and although it measures continuously for a certain time, it is not reversible. Accordingly, in this piezoelectrical system, whenever a new sample has to be measured the chemical sensing tip has to be restored to activity by washing at pH 3 at which the conjugates are dissociated.
It is an object of the present invention to provide a new concept for chemical and biochemical sensors which combines effectively the features of stability, specificity and reversibility without restriction on the size of the analyte.
Another object of the invention is to provide a sensor which does not need a semipermeable membrane while still keeping a constant predetermined amount of analogue receptors or analogue ligands (labelled or unlabelled) in the sensing area. This type of sensor is more suitable for in-vivo applications.
The new concept is also applicable to a probe as well as for immunoassay without the need to add extra reagents, for example an analog-analyte as normally used in the conventional immunoassays.